EP3675959A1 - Ensemble transducteur pour générer des ultrasons focalisés - Google Patents

Ensemble transducteur pour générer des ultrasons focalisés

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Publication number
EP3675959A1
EP3675959A1 EP18849828.1A EP18849828A EP3675959A1 EP 3675959 A1 EP3675959 A1 EP 3675959A1 EP 18849828 A EP18849828 A EP 18849828A EP 3675959 A1 EP3675959 A1 EP 3675959A1
Authority
EP
European Patent Office
Prior art keywords
acoustic
transducer assembly
ultrasound transducer
layer
ultrasound
Prior art date
Legal status (The legal status is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the status listed.)
Pending
Application number
EP18849828.1A
Other languages
German (de)
English (en)
Other versions
EP3675959A4 (fr
Inventor
Jeffrey Kyle WOODACRE
Jeremy Brown
Current Assignee (The listed assignees may be inaccurate. Google has not performed a legal analysis and makes no representation or warranty as to the accuracy of the list.)
Dalhousie University
Original Assignee
Dalhousie University
Priority date (The priority date is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the date listed.)
Filing date
Publication date
Application filed by Dalhousie University filed Critical Dalhousie University
Publication of EP3675959A1 publication Critical patent/EP3675959A1/fr
Publication of EP3675959A4 publication Critical patent/EP3675959A4/fr
Pending legal-status Critical Current

Links

Classifications

    • B—PERFORMING OPERATIONS; TRANSPORTING
    • B06—GENERATING OR TRANSMITTING MECHANICAL VIBRATIONS IN GENERAL
    • B06B—METHODS OR APPARATUS FOR GENERATING OR TRANSMITTING MECHANICAL VIBRATIONS OF INFRASONIC, SONIC, OR ULTRASONIC FREQUENCY, e.g. FOR PERFORMING MECHANICAL WORK IN GENERAL
    • B06B1/00—Methods or apparatus for generating mechanical vibrations of infrasonic, sonic, or ultrasonic frequency
    • B06B1/02—Methods or apparatus for generating mechanical vibrations of infrasonic, sonic, or ultrasonic frequency making use of electrical energy
    • B06B1/06—Methods or apparatus for generating mechanical vibrations of infrasonic, sonic, or ultrasonic frequency making use of electrical energy operating with piezoelectric effect or with electrostriction
    • B06B1/0644—Methods or apparatus for generating mechanical vibrations of infrasonic, sonic, or ultrasonic frequency making use of electrical energy operating with piezoelectric effect or with electrostriction using a single piezoelectric element
    • B06B1/0662—Methods or apparatus for generating mechanical vibrations of infrasonic, sonic, or ultrasonic frequency making use of electrical energy operating with piezoelectric effect or with electrostriction using a single piezoelectric element with an electrode on the sensitive surface
    • B06B1/067—Methods or apparatus for generating mechanical vibrations of infrasonic, sonic, or ultrasonic frequency making use of electrical energy operating with piezoelectric effect or with electrostriction using a single piezoelectric element with an electrode on the sensitive surface which is used as, or combined with, an impedance matching layer
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61N—ELECTROTHERAPY; MAGNETOTHERAPY; RADIATION THERAPY; ULTRASOUND THERAPY
    • A61N7/00—Ultrasound therapy
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61N—ELECTROTHERAPY; MAGNETOTHERAPY; RADIATION THERAPY; ULTRASOUND THERAPY
    • A61N7/00—Ultrasound therapy
    • A61N7/02—Localised ultrasound hyperthermia
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B17/00—Surgical instruments, devices or methods, e.g. tourniquets
    • A61B17/22—Implements for squeezing-off ulcers or the like on the inside of inner organs of the body; Implements for scraping-out cavities of body organs, e.g. bones; Calculus removers; Calculus smashing apparatus; Apparatus for removing obstructions in blood vessels, not otherwise provided for
    • A61B17/225—Implements for squeezing-off ulcers or the like on the inside of inner organs of the body; Implements for scraping-out cavities of body organs, e.g. bones; Calculus removers; Calculus smashing apparatus; Apparatus for removing obstructions in blood vessels, not otherwise provided for for extracorporeal shock wave lithotripsy [ESWL], e.g. by using ultrasonic waves
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61N—ELECTROTHERAPY; MAGNETOTHERAPY; RADIATION THERAPY; ULTRASOUND THERAPY
    • A61N7/00—Ultrasound therapy
    • A61N2007/0004—Applications of ultrasound therapy
    • A61N2007/0021—Neural system treatment
    • A61N2007/003—Destruction of nerve tissue
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61N—ELECTROTHERAPY; MAGNETOTHERAPY; RADIATION THERAPY; ULTRASOUND THERAPY
    • A61N7/00—Ultrasound therapy
    • A61N2007/0039—Ultrasound therapy using microbubbles
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61N—ELECTROTHERAPY; MAGNETOTHERAPY; RADIATION THERAPY; ULTRASOUND THERAPY
    • A61N7/00—Ultrasound therapy
    • A61N2007/0056—Beam shaping elements
    • A61N2007/006—Lenses
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61N—ELECTROTHERAPY; MAGNETOTHERAPY; RADIATION THERAPY; ULTRASOUND THERAPY
    • A61N7/00—Ultrasound therapy
    • A61N2007/0056—Beam shaping elements
    • A61N2007/0065—Concave transducers
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61N—ELECTROTHERAPY; MAGNETOTHERAPY; RADIATION THERAPY; ULTRASOUND THERAPY
    • A61N7/00—Ultrasound therapy
    • A61N2007/0082—Scanning transducers

Definitions

  • the present disclosure relates to focused ultrasound.
  • the present disclosure relates to focused ultrasound for therapeutic applications, such as histotripsy and high intensity focused ultrasound (HIFU).
  • HIFU high intensity focused ultrasound
  • Histotripsy is a tissue ablation process in which short bursts of high intensity ultrasound are focused to a small focal region, exceeding the vapor pressure within the focal region and causing gas bubbles to form. When these bubbles collapse, the Shockwave destroys the tissue structure, leaving liquified remnants of the original tissue at the focal region.
  • HIFU High-intensity focused ultrasound
  • heat can also damage the surrounding tissue.
  • Both HIFU and histotripsy transducers function similarly, as both require the focusing of ultrasound to a small focal region.
  • the piezoelectric layer is a composite piezoelectric material having an acoustic impedance configured to match the acoustic impedance of the acoustic lens.
  • the acoustic lens may be formed from aluminum, or an alloy thereof, and may have a distal surface having a non-spherical profile for producing a focal region that is smaller than an equivalent spherical lens.
  • the acoustic lens may have an f-number less than two.
  • the acoustic lens is coated with a polymer acoustic impedance matching layer that is compatible with deposition via chemical vapor deposition, such as a p-xylylene based polymer.
  • a polymer acoustic impedance matching layer that is compatible with deposition via chemical vapor deposition, such as a p-xylylene based polymer.
  • the acoustic lens is formed from aluminum or an alloy thereof, and the polymer acoustic impedance matching layer is a Parylene layer.
  • an ultrasound transducer assembly for generating focused ultrasound, the ultrasound transducer assembly comprising:
  • an acoustic lens having a proximal surface and a curved distal surface, wherein the proximal surface is attached to the piezoelectric layer;
  • an acoustic impedance of the piezoelectric layer approximately matches an acoustic impedance of the acoustic lens.
  • an ultrasound transducer assembly for generating focused ultrasound, the ultrasound transducer assembly comprising: a composite piezoelectric layer;
  • an acoustic lens having a proximal surface and a curved distal surface, wherein the proximal surface is attached to the composite piezoelectric layer, wherein the acoustic lens comprises 85% aluminum by weight;
  • a polymer acoustic impedance matching layer coating the curved distal surface of the acoustic lens, wherein the polymer acoustic impedance matching layer is formed from a p-xylylene based polymer;
  • an acoustic impedance of the composite piezoelectric layer matches an acoustic impedance of the acoustic lens within +- 40%
  • curved distal surface has an elliptical shape
  • the acoustic lens has an f-number of less than two.
  • an ultrasound system for generating focused ultrasound comprising:
  • an ultrasound transducer assembly comprising:
  • an acoustic lens having a proximal surface and a curved distal surface, wherein the proximal surface is attached to the piezoelectric layer; and an acoustic impedance matching layer coating the curved distal surface of the acoustic lens;
  • driver circuitry operably connected with the ultrasound transducer assembly, wherein the driver circuitry is configured to deliver electrical pulses with a voltage and operating frequency sufficient for generating ultrasound pulses for performing histotripsy;
  • an acoustic impedance of the piezoelectric layer approximately matches an acoustic impedance of the acoustic lens
  • the acoustic impedance matching layer is a quarter wave matching layer.
  • FIG. 1 shows an example transducer assembly for generating therapeutic focused ultrasound, in which a piezoelectric layer is coupled to an acoustic lens.
  • FIG. 2 shows an example transducer assembly in which a piezoelectric layer is coupled to an acoustic lens for generating and therapeutic focused ultrasound, and where the transducer assembly includes a coaxial imaging transducer for co- registered imaging and ultrasound therapy.
  • FIG. 3 shows an example system for generating therapeutic focused ultrasound.
  • FIG. 4 shows an example system for generating therapeutic focused ultrasound and performing ultrasound imaging.
  • FIG. 5 is a photograph of an example focused ultrasound assembly including a piezoelectric composite that adhered to and impedance matched to an elliptical aluminum acoustic lens (6061 aluminum), where the lens is coated with a layer of Parylene C.
  • FIGS. 6A and 6B plot, respectively, the measured electrical impedance and phase of an example transducer composite without a lens attached, shown results from simulations employing a KLM model, illustrating the close correspondence between the measured and simulated results.
  • the resonance peak from FIGS. 6A and 6B are observed to spread over the spectrum due to the presence of the lens.
  • FIGS. 7A and 7B plot impedance and phase measurements for a composite transducer assembly having an integrated acoustic lens.
  • FIG. 7C shows the electrical impedance magnitude of composite-lens stack along with KLM models both with and without the bonding epoxy between composite and lens. Modelling the bonding epoxy is necessary to ensure the KLM resonance matches the measured resonance. Discrepancies are likely due to the inability of KLM to model internal lens reflections.
  • FIG. 7D shows the electrical impedance phase where, again, the addition of 20 ⁇ of epoxy allows KLM to more closely match the phase characteristics.
  • FIGS. 8A and 8B plot KLM model impedance spectra for transducers with 0,
  • FIG. 9 plots the power spectrum output for each epoxy layer thickness, showing the power output doubling from the case of a device without epoxy to a device having a 120 ⁇ layer of epoxy. As can be seen in the figure, the bandwidth is reduced as the epoxy thickness is increased, as indicated by the much narrower peak for the 120 ⁇ epoxy layer compared to the broadness of the "no epoxy" line.
  • FIG. 10 plots simulated transducer efficiency for a device without epoxy as well as devices with 20, 60, 120, and 160 micrometers of epoxy, as a function of drive frequency. The addition of epoxy increases efficiency, but narrows the band in which the system is more efficient.
  • FIGS. 11A and 11B plot the measured pressure profile for the transducer assembly without an integrated co-registered imaging transducer, when the therapeutic transducer is driven at a drive voltage of 20 V, a frequency of 6.8 MHz, and 20 cycles in the burst signal.
  • FIGS. 12A and 12B plot the peak-to-peak pressure vs. drive voltage for the non-imaging transducer.
  • FIGS. 13A and 13B are photographs showing a bubble cloud.
  • the photograph shown in FIG. 10A was generated with a 170V, 3 cycle pulse with a 10 ms repetition rate.
  • the photograph shown in FIG. 10B was generated in degassed, deionized water with a 6.8 MHz single-cycle, single-ended 173 V pulse at a 50 Hz repetition rate.
  • the needle tip shown the right is a 26 gauge needle with a diameter of 0.46 mm, demonstrating that the cloud size measures ⁇ 0.2 mm diameter vertically at its smallest size.
  • FIG. 14 is a photograph of an example imaging and ablation device, where the histotripsy ablation lens has a center-hole to allow an imaging tool to visualize the area in real-time during ablation.
  • This example imaging device employs a 64- element, 40 MHz phased array endoscope.
  • FIG. 15A and 15B plot the peak-to-peak pressure vs. drive voltage for a transducer assembly having an integrated and co-registered imaging transducer. A one-way single-cycle pulse response is also shown in the top-left corner of FIG. 12A.
  • FIGS. 16A and 16B plot the measured pressure profile for the transducer assembly with an integrated co-registered imaging transducer, when the therapeutic transducer is driven at a drive voltage of 40 V, a frequency of 6.8 MHz, and 20 cycles in the burst signal.
  • FIGS. 17A and 17B show (A) a chinchilla cerebellum imaged showing the molecular layer (dark layer), the granular layer (highly specular), and white matter tracts (thin dark lines in granular layer) and (B) a histotripsy bubble cloud, visible as a highly specular region near the image center, has been plunged into the cerebellum. Both images were collected in real-time using a co-registered 40 MHz endoscopic phased array.
  • the terms “comprises” and “comprising” are to be construed as being inclusive and open ended, and not exclusive. Specifically, when used in the specification and claims, the terms “comprises” and “comprising” and variations thereof mean the specified features, steps or components are included. These terms are not to be interpreted to exclude the presence of other features, steps or components.
  • exemplary means “serving as an example, instance, or illustration,” and should not be construed as preferred or advantageous over other configurations disclosed herein.
  • the terms “about” and “approximately” are meant to cover variations that may exist in the upper and lower limits of the ranges of values, such as variations in properties, parameters, and dimensions. Unless otherwise specified, the terms “about” and “approximately” mean plus or minus 25 percent or less.
  • any specified range or group is as a shorthand way of referring to each and every member of a range or group individually, as well as each and every possible sub-range or sub -group encompassed therein and similarly with respect to any sub-ranges or sub-groups therein. Unless otherwise specified, the present disclosure relates to and explicitly incorporates each and every specific member and combination of sub-ranges or subgroups.
  • the term "on the order of”, when used in conjunction with a quantity or parameter, refers to a range spanning approximately one tenth to ten times the stated quantity or parameter.
  • the phrase "co-registered" when employed with reference to the relationship between a therapeutic ultrasound transducer and an imaging ultrasound transducer refers to a fixed mechanical relationship between the therapeutic ultrasound transducer and the imaging ultrasound transducer, where a focal region of the therapeutic ultrasound transducer lies within an imaging region of the imaging ultrasound transducer.
  • Various embodiments of the present disclosure provide systems and devices for generating therapeutic focused ultrasound, for example, for therapeutic applications such as tissue ablation.
  • the example embodiments described herein arose from technical solutions that were discovered when addressing technical problems encountered when attempting to fabricated curved piezoelectric materials for use in histotripsy applications.
  • the present inventors initially fabricated histotripsy transducers from a 1-3 connected composite (made of piezoelectric pillars interspersed in a matrix of epoxy).
  • Soft epoxies were employed, permitting the bending of the composite to facilitate focusing to some degree.
  • soft epoxies it was found that soft epoxies curved well, but failed to hold their shape and were not capable of providing an appropriate composite thickness to achieve a desired control over transducer frequency.
  • the present inventors seeking a solution to this problem, employed an alternative design involving the use of an acoustic lens formed in a passive material that is attached to an active piezoelectric composite substrate.
  • An example of such an embodiment is illustrated in FIG. 1.
  • the transducer assembly includes a piezoelectric layer 100 that is attached to an acoustic lens 1 10 at a proximal surface 1 12 thereof.
  • the distal surface 1 14 of the acoustic lens is coated with a polymer impedance matching layer 120 having a thickness (e.g.
  • acoustic impedance suitable for impedance matching between the acoustic impedance of the acoustic lens 1 10 and the acoustic impedance of tissue e.g. the acoustic impedance of water.
  • the distal surface of the piezoelectric layer 100 may be adhered to the proximal surface 1 12 of the acoustic lens 1 10 using a thin layer of adhesive, such as a thin layer of epoxy.
  • the thickness of the epoxy may be selected to be less than the wavelength associated with the operating frequency of the ultrasound transducer assembly to ensure good transmission acoustic energy into the acoustic lens 1 10.
  • an intermediate layer may be provided between the acoustic lens and the piezoelectric layer, where the acoustic impedance of the intermediate layer is lower than the respective acoustic impedances of the acoustic lens and the piezoelectric layer, and wherein the thickness of the intermediate layer is selected to control or achieve one or more device performance parameters, such as, but not limited to, power transfer, emission bandwidth, and resonant frequency.
  • the thickness of the intermediate layer may be selected to modify or select the transducer output power for a given input power, where, for example, increasing film thickness may provide improved acoustic power output for a given input voltage relative to a case with a thin ( ⁇ 20 ⁇ ) intermediate layer.
  • the intermediate layer may be electrically conductive or non-conductive.
  • the intermediate layer may have a thickness between 15 and 25 microns, or a thickness between 15 and 50 microns, or a thickness between 20 and 50 microns, or a thickness between 25 and 50 microns, or a thickness between 50 and 70 microns, or a thickness between 50 and 100 microns, or a thickness between 50 and 150 microns, or a thickness between 100 and 150 microns, or a thickness between 100 and 200 microns, or a thickness of at least 15 microns, or a thickness of at least 20 microns, or a thickness of at least 50 microns, or a thickness of at least 75 microns, or a thickness of at least 100 microns.
  • the intermediate layer may be formed from a wide variety of materials.
  • the intermediate layer may be formed from an adhesive, such as an epoxy or glue.
  • the intermediate layer may be a material other than an adhesive, such as a liquid layer, for example, oil or water.
  • the matching layer 120 may be a quarter wave matching layer corresponding to the frequency associated with the peak in the acoustic power spectrum of the transducer assembly, where the frequency of the peak is dependent on the thickness of the intermediate layer.
  • the drive frequency of the transducer assembly may correspond to (e.g. coincide with, or lie within a 3 dB bandwidth) the frequency associated with a peak in the acoustic power spectrum of the transducer assembly, where the frequency of the peak is dependent on the thickness of the intermediate layer.
  • the thickness of the intermediate layer may be selected to be sufficient to effect an increase in peak emitted acoustic power in the acoustic power spectrum of the ultrasound transducer assembly by at least two or three relative to an equivalent ultrasound transducer assembly absent of the intermediate layer (as may be determined, for example, via experimentation and/or simulation). In some example implementations, the thickness of the intermediate layer may be selected to be sufficient to effect an increase in peak efficiency in an acoustic efficiency spectrum of the ultrasound transducer assembly by at least 20% or 40% relative to an equivalent ultrasound transducer assembly absent of the intermediate layer.
  • an intermediate layer may be provided between a piezoelectric material and a second material to increase the overall power transfer from the piezoelectric to the second material for a given drive voltage of the piezoelectric, while additionally shifting the drive frequency at which maximum output power occurs, where the magnitude of this shift depends on layer thickness, where the acoustic impedance of the intermediate layer is lower than the acoustic impedances of the piezoelectric material and the second material.
  • FIG. 1 shows an example of a non- limiting implementation of a housing 130 for mechanically supporting the
  • the acoustic lens 1 10 and the piezoelectric layer 100 may be recessed within a housing 130 that is non-conductive.
  • Electrical cabling 125 e.g. a coaxial cable
  • enters the housing 130 through a hole or aperture 135 e.g. which may be sealed with a water- resistant epoxy, adhesive or other sealing material, and electrical connections are made within the housing 130 between the drive and ground wires and respective surfaces of the piezoelectric layer 100.
  • air backing is employed to ensure that the acoustic energy emerges from the front of the device and into the tissue.
  • other backing materials could be used in the alternative, albeit with a reduction in the device output.
  • FIG. 1 illustrates a non-limiting example embodiment in which the acoustic lens 1 10 is electrically conductive, where the ground connection to the distal surface of the piezoelectric layer 100 is made indirectly through the conductive acoustic lens 1 10, and through a conductive adhesive layer that attaches the piezoelectric layer 100 to the acoustic lens 1 10.
  • the ground connection may be connected to the acoustic lens 1 10, for example, via conductive epoxy or simply via contacted under the application of pressure (e.g. using a spring or set screw; not shown).
  • the present inventors found that in order to achieve sufficient focusing for generating and sustaining a bubble cloud for histotripsy applications, the conventional approach involving a circular lens shape was insufficient, as spherical aberrations hampered the ability to achieve a sufficiently small focus.
  • the present inventors adapted the design such that the outer curved surface of the acoustic lens is non-spherical, thereby avoiding spherical aberrations and facilitating a tighter focus than that achievable with a lens having a spherical surface.
  • an elliptical lens surface was selected in order to achieve improved focusing.
  • the inventors found that it is beneficial for the acoustic lens to have a low f-number, such as an f-number of less than one (unity), or an f-number less than two, in order to facilitate the generation of a sufficiently strong focus for histotripsy applications.
  • the inventors selected aluminum (or alloys thereof, e.g. a metal alloy containing at least 85% aluminum by weight) as a suitable material for the acoustic lens, since aluminum is a low-cost material that could be readily machined (e.g. using CNC machining) in order to achieve the high curvature needed to a produce a low-f-number acoustic lens, and to achieve the non-spherical (e.g.
  • the elliptical lens shape avoids spherical aberrations in the lens, which are important to avoid when the lens focus is smaller than the lens aperture.
  • the inventors Having selected a suitable material for machining the desired curvature of the acoustic lens, the inventors then considered the acoustic impedances of the piezoelectric layer 100, the acoustic lens 1 10, and the polymer acoustic impedance matching layer 120.
  • Aluminum has an acoustic impedance of approximately 17 MRayls, which is significantly lower than conventional piezoelectric materials.
  • the piezoelectric layer 100 was therefore configured as a piezoelectric composite with a volume fraction of piezoelectric material selected to achieve or approximate acoustic impedance matching between the piezoelectric layer 100 and the acoustic lens 1 10.
  • the acoustic impedance of the piezoelectric layer approximately matches the acoustic impedance of the acoustic lens when the acoustic impedance of the piezoelectric layer is within ⁇ 40% of the acoustic impedance of the acoustic lens.
  • a composite piezoelectric layer may be formed as a 1-3 piezoelectric composite having a volume fraction and pillar geometry suitable to achieve or approximate the impedance matching condition, with an acoustic impedance equal to, or approximately equal to, 17 MRayls.
  • the acoustic impedance of the composite piezoelectric layer may lie between 16 and 18 MRayls, between 15 and 19 MRayls, between 14 and 20 MRayls, or between 13 and 21 MRayls.
  • Such an acoustic impedance matches (or approximately matches) the acoustic impedance of the aluminum lens, ensuring that a substantial fraction of the acoustic energy outputted by the composite is transferred into the acoustic lens.
  • the piezoelectric composite need not have a 1-3 configuration, and other composites, such as a 2-2 composite, may be employed in the alternative. While some of the example implementations described herein involve the use of a composite piezoelectric layer, it will be understood that other example implementations may involve non-composite piezoelectric layers.
  • the composite could be made from any form of piezoelectric ceramic (e.g. PZT-4, PZT-5A, PZT-5H, PMN-PT), although the choice of ceramic will affect the drive voltage necessary to perform histotripsy as well as the saturation voltage at which driving the piezoelectric harder no longer increases pressure output.
  • piezoelectric ceramic e.g. PZT-4, PZT-5A, PZT-5H, PMN-PT
  • Single-crystal PMN-PT piezoelectric could also be used.
  • Parylene C was selected as a suitable polymer for forming the polymer impedance matching layer 120.
  • the acoustic impedance of Parylene C is approximately 2.7 MRayl, this acoustic impedance is close to the target acoustic impedance that is predicted by the following acoustic impedance matching equation for a single ⁇ /4 impedance matching layer:
  • Z? is the impedance of aluminum ( ⁇ 17 MRayls) and Z 2 is the acoustic impedance of water ( ⁇ 1.5 MRayls).
  • Z? is the impedance of aluminum ( ⁇ 17 MRayls) and Z 2 is the acoustic impedance of water ( ⁇ 1.5 MRayls).
  • other types of p-xylylene based polymers may alternatively be employed, such as Parylene N or Parylene D.
  • an p-xylylene based polymer for the acoustic impedance matching layer 120 is beneficial in that it is compatible with chemical vapor deposition (CVD).
  • CVD is particularly beneficial in the context of example embodiments involving the formation of an acoustic impedance matching layer on the highly curved surface of low (e.g. sub-unity, or less than two) f-number acoustic lens, because CVD can achieve a uniform coating thickness even in the presence of high curvature.
  • the polymer impedance matching layer is formed from Parylene C
  • other types of polymers may be used.
  • CVD-compatible polymers having acoustic impedances in the range of 2.5-4 MRayls, may be alternatively employed, such as polyimide or fluoropolymers such as Teflon.
  • a glass-based material may alternatively be employed for the acoustic lens in order to achieve a suitable set of materials for acoustic impedance matching or approximate impedance matching.
  • glasses such as quartz glass and silica glass have acoustic impedance values in the range of 13-15 MRayls.
  • f-number such as, in one example embodiment, an f-number less than unity, or in another example embodiment, an f-number less than two.
  • Beam width and depth-of-field for the therapeutic focused transducer are linearly and quadratically proportional to f-number (F#) respectively, so increasing F# causes the size of the focus and, therefore, the volume over which the ultrasound energy is spread, to increase, requiring a higher drive voltage to the transducer to reach cavitation pressures. Also, since the focus is less tight with increasing F#, the ability to target ablation to specific areas diminishes.
  • a low F# lens such as an F# less than unity, or an F# less than two
  • a non-spherical lens such as an elliptical lens
  • spherical aberrations would otherwise result in a larger effective focal area, and would also reduce the pressure within the focal region and require higher drive voltages.
  • the transducer assembly includes a co-registered imaging transducer.
  • the imaging transducer 140 is received and housed in a coaxial manner relative to the acoustic lens 1 10, with the distal end of the imaging transducer emerging through an aperture formed in the distal surface of the acoustic lens 1 10.
  • the imaging transducer is a miniaturized endoscopic transducer, e.g. having a diameter less than 4 mm, permitting the housing of the imaging transducer within an acoustic lens having a diameter less than 10 mm.
  • the imaging transducer may be supported within the transducer assembly according to a wide range of example implementations, and a specific method may depend on the device geometry.
  • set screws may be provided in the housing that gently press into the phased array device, supporting it in place, in a removable fashion.
  • the imaging transducer 140 may be permanently adhered within the transducer assembly, such as via an adhesive.
  • Non-limiting examples of the ultrasound imaging transducer include the ultrasound endoscope described in US Patent Publication No. 2015/0209005A1 (Bezanson et al.), titled "Ultrasound Endoscope and Methods of Manufacture
  • WO2017127328 A1 (Brown et al.), titled “Compact Ultrasound Device Having Annular Ultrasound Array Peripherally Electrically Connected to Flexible Printed Circuit Board and Method of Assembly Thereof", which is incorporated herein by reference in its entirety. It will be understood that these devices are provided only as illustrative examples, and that a wide variety of ultrasound imaging devices may be integrated within the ultrasound transducer assembly.
  • ultrasound imaging transducers that are disposed co-axially with the axis of the acoustic lens of the therapeutic transducer
  • the imaging axis need not be co-axial with the axis of the acoustic lens, provided that the focal region of the therapeutic ultrasound lies within an imaging region associated with the imaging transducer.
  • the configuration shown in FIG. 2 provides one example implementation of the orientation of the ablation lens relative to the endoscope.
  • the endoscope tip may be recessed from the lens curvature so that the endoscope does not occlude the ablation tool, while still having the ablation zone centered in the imaging window.
  • the lens-composite stack may be encased to ensure the composite remains air-backed.
  • FIG. 3 provides a block diagram illustrating an example implementation of a system for performing procedures involving focused ultrasound, such as histotripsy or HIFU.
  • the control hardware 200 is operably connected to the transducer driver electronics/circuitry 260, which drives the transducer assembly 270 to generate focused ultrasound.
  • the driver electronics/circuitry 260 may comprise, for example, a high-voltage single ended ultrasound pulser.
  • the driver in the case of a histotripsy system, the driver
  • electronics/circuitry 260 may be capable of supplying over 200 volts (e.g. over 400, 600, 800 or 900 V), for example, at up to 144 amps of pulsed current, or 24 amps continuous current.
  • the control hardware 200 includes one or more processors 210 (for example, a CPU/microprocessor), bus 205, memory 215, which may include random access memory (RAM) and/or read only memory (ROM), a data acquisition interface 220, a display 225, external storage 230, one more communications interfaces 235, a power supply 240, and one or more input/output devices and/or interfaces 245 (e.g. a speaker, a user input device, such as a keyboard, a keypad, a mouse, a position tracked stylus, a position tracked probe, a foot switch, and/or a microphone for capturing speech commands).
  • processors 210 for example, a CPU/microprocessor
  • bus 205 memory 215, which may include random access memory (RAM) and/or read only memory (ROM), a data acquisition interface 220, a display 225, external storage 230, one more communications interfaces 235, a power supply 240, and one or more input/output devices and/or interfaces 245 (e.
  • the control hardware 200 may be programmed with programs, subroutines, applications or modules, such as transducer control module 255, which include executable instructions, which when executed by the one or more processors 210, causes the system to generate a series of pulses suitable for a selected type of focused ultrasound therapy. Such instructions may be stored, for example, in memory 215 and/or other storage.
  • control hardware 200 may be implemented as one or more physical devices that are coupled to processor 210 through one of more communications channels or interfaces.
  • control hardware 200 can be implemented using application specific integrated circuits (ASICs).
  • ASICs application specific integrated circuits
  • control hardware 200 can be implemented as a combination of hardware and software, where the software is loaded into the processor from the memory or over a network connection.
  • FIG. 4 illustrates an example embodiment of a system configured for focused ultrasound therapy and imaging using a transducer assembly 272 having an integrated co-registered imaging transducer and a focused ultrasound therapeutic transducer.
  • Control hardware 200 is employed to control transmit beamformer 282 and receive beamformer 284, and Tx/Rx switch 285, and for processing the beamformed receive signals.
  • control hardware 200 may include an image processing module 256 for processing image data obtained using the co-registered imaging transducer.
  • the example embodiments disclosed above may be employed for a wide variety of applications, such as neurological procedures.
  • one of the most common methods of tumor resection in the brain is the use of burr-hole surgery, in which a hole is made in the skull and, using visual guidance as well as pre-operative MRI images, a cavitational ultrasonic surgical aspiration (CUSA) device is used to ablate the tumor tissue.
  • This CUSA treatment is a contact ablation where the device cavitates and ablates tissue adjacent to the tip and the liquified tissue is then pulled away via suction.
  • the systems and devices disclosed herein may be employed for focused ultrasound applications other than histotripsy.
  • the systems and devices disclosed herein can be used for HIFU.
  • the embodiments disclosed herein could be used for HIFU without the presence of the polymer acoustic impedance matching layer, and/or with a different lens material, and/or using a non-composite
  • the distal lens shape may be spherical in the case of HIFU applications, as HIFU does not typically require the high intensity of ultrasound that histotripsy does, and instead involves the sustained delivery of ultrasound to heat tissue.
  • a small, 10 mm aperture histotripsy transducer was characterized and, following characterization, the device was modified to include a co-linear, co-registered 40 MHz imaging device to allow both imaging and ablation in real-time with a 10 mm aperture.
  • the ultrasound ablation transducer described in the present example consists of an air-backed piezoelectric composite bound to an aluminum lens using epoxy, where the aluminum lens has a quarter wavelength matching layer on the front face matching to water.
  • a photograph of the assembled composite transducer and coated parabolic aluminum acoustic lens is shown in FIG. 5.
  • the piezoelectric composite was designed to provide maximum
  • the aluminum lens was fabricated using a CNC milling process and designed to focus at a 7 mm depth.
  • the curvature of the lens was fabricated to be elliptical, as per the equation provided above.
  • elliptical lens shape avoids spherical aberrations in the lens, which are important to avoid when the lens focus is smaller than the lens aperture.
  • Parylene has a longitudinal speed of sound of 2135 m/s with an acoustic impedance of 2.75 MRayls [20], so a layer
  • the epoxy film thickness and properties could be chosen to modify the transducer output, where, for example, increasing film thickness attains improved acoustic power output for a given input voltage but with a reduced efficiency and bandwidth, as well as a frequency shift of the resonance.
  • KLM models show the film thickness is near 20 ⁇ , approximately doubling output power compared to having no film for a fixed drive-voltage while also shifting the maximum output frequency from 5 MHz to 6.8 MHz. This doubling of output power should result in a sqrt(2) gain in acoustic pressure at the lens focus.
  • Characterization of the transducer was performed by the following methods: electrical impedance measurements both before and after lens bonding, measuring a pressure field map near the transducer focus, obtaining a measure of the peak minimum focal pressure versus drive voltage, and imaging of a single-cycle pulse generated bubble cloud.
  • the experimental electrical impedance measured using an Agilent 4294A Precision Impedance Analyzer (Agilent Technologies, Santa Clara, USA), and the KLM derived electrical impedance are shown in FIGS. 6A and 6B for a 10 mm diameter disc of air backed and air loaded composite prior to attaching a lens.
  • KLM code was used to model the transducer electrical impedance. This code was able to model not only a bare piezoelectric composite, but the full transducer stack which included an epoxy bonding layer and aluminum lens. To include the lens in the KLM model, the composite-lens stack was divided into a set of concentric rings of equal area, each with equal backing, composite, epoxy and Parylene layers, but with aluminum layer of varying height to account for the lens curvature varying as a function of radius.
  • a ring would be modeled as having a 2 mm aluminum layer as part of the lens, while a ring at the outer edge would be modeled with a 4.4 mm aluminum layer as the lens curvature is larger at the outer edge.
  • FIGS. 7A and 7B show the measured impedance and phase curves, respectively, for a transducer with the lens attached.
  • the resonance peak normally seen for a composite is absent, as the lens spreads the resonances out over a wide band.
  • near 8 MHz in the measured impedance is a second resonance feature corresponding to a lateral mode of the composite which is not captured in KLM, as KLM is inherently a one-dimensional model. This one-dimensional limitation also leads to difficulties in modeling the full transducer electrical impedance when a lens is introduced.
  • FIGS. 7C and 7D show a comparison is shown between the measured electrical impedance for a composite with lens attached, along with KLM models both with and without Epotek 301 between the composite and lens.
  • the measured impedance magnitude in FIG. 7C shows a resonance at 5 MHz and an anti-resonance at 7 MHz; however, a KLM model with composite in direct contact to the lens shows only a decreasing impedance magnitude with increasing frequency.
  • the addition of a 20 ⁇ epoxy layer is therefore necessary to capture the measured system behavior in the model.
  • Also of note are a number of small, 5 Ohm amplitude oscillatory features between 2 MHz and 7 MHz that are visible in the measured impedance.
  • the impedance phase curve in FIG. 7D also shows the importance of an epoxy layer where the 20 ⁇ epoxy layer moves the maximum phase point from 4 MHz to 6.5 MHz and increasing the value from -60 degrees to -20 degrees, again, more closely matching to the measured electrical impedance.
  • FIGS. 8A and 8B the simulated electrical impedance of five transducer designs is shown, with each device having an increasing epoxy layer thickness.
  • the resonance minimum point
  • the anti-resonance peak
  • the phase peak increases while the frequency also shifts to the left.
  • the shift in frequency may be, for example, mass-spring related, or may be a resonance cavity created by the addition of the epoxy.
  • FIG. 9 plots of the simulated acoustic power output normalized to input voltage squared show a power increase with increasing epoxy layer thickness, up to a limit.
  • a quarter wavelength matching layer is added to the front transducer face with a thickness corresponding to wavelength associated with the maximum transducer output power prior to the quarter wavelength layer being added; or, in other words, the quarter wavelength layer is matched to work in tandem with the specific epoxy layer thickness.
  • the simulated devices having layer thickness of 60 micrometers and 120 micrometers of provide an almost identical power output, with a shifted frequency between them, while increasing further to a 640 micrometer layer creates a split output with reduced power in each band, compared to the devices with 60 and 120 micrometer layers.
  • power output is increased by a factor of up to 3.75 by adding the epoxy layer.
  • FIG. 10 illustrates how simulated efficiency changes as the bonding epoxy layer is widened, where it is observed that the epoxy layer increases efficiency, albeit over a narrower band of frequencies.
  • FIG. 1 1 A plots the 2D pressure profile where the - 6dB radial beam width is 0.207 mm, and the -6dB axial beam length is 1.061 mm.
  • the pressure field was measured in a 5 mm x 2 mm plane intersecting the transducer focus where the plane normal vector is perpendicular to the direction of acoustic propagation.
  • the 0.04 mm needle hydrophone was scanned through the pressure field, where an oscilloscope recorded the peak-negative pressure while the transducer was driven using a pulser based on the design by Brown and Lockwood [21] with a 20 cycle, 20 V, 6.8 MHz pulse.
  • FIG. 11 B shows that the -3 dB width in the radial direction is 0.145 mm while in the axial direction the -3 dB width measures 0.698 mm.
  • FIG. 12A the peak-to-peak pressure versus drive voltage for the non- imaging transducer is shown.
  • the measurement was limited to below 25 V as, above this, large oscillations in the hydrophone measurements were observed, which were believed to be the beginnings of cavitation, which could potentially damage the hydrophone.
  • FIG. 12B shows an alternative representation of the data plotted in FIG. 12A, plotting the peak-negative pressure (the maximum negative value of pressure instead of the difference between the maximum and minimum pressure).
  • the inset to FIG. 12B plots a representative one-way single-cycle pulse response as measured at the hydrophone to provide device bandwidth. Within the inset, temporal ringing following the main pulse is likely a result of reverberation within the lens, and potentially also ringing within the hydrophone which could not be uncoupled from the measurement.
  • a linear relationship between pressure and drive voltage is seen from 2.5 V up to 17 V. Above 17 V, initial evidence of cavitation at the hydrophone tip was seen as noise in the oscilloscope signal. At 25 V, cavitation at the hydrophone tip made pressure measurements inconsistent and, therefore, measurements were stopped.
  • the inset plot shows a one-way, single-cycle pulse response as measured with the hydrophone. The ringing after the initial pulse is due to reverberations within the lens. The pulse bandwidth was 59%.
  • the transducer was driven with a 20-cycle pulse train to ensure steady-state was reached at 6.8 MHz with a pulse repetition frequency of 100 Hz.
  • a minimum drive voltage of 2.3 V was needed for the drive circuit to power the transducer, after which point the pressure follows a linear trend of 0.29 MPa / V up to 25 V.
  • a drive voltage of greater than 90 V should be needed for a multi-cycle pulse.
  • the bubble cloud was identified visibly using a microscope, where it can be seen in FIG. 13A at 300.
  • the bubble cloud was generated with a 170V, 3 cycle pulse with a 10 ms repetition rate.
  • the needle tip shown the right (at 310) is a 26 gauge needle with a diameter of 0.46 mm, demonstrating that the cloud size measures ⁇ 0.2 mm diameter vertically at its smallest size.
  • the bubble cloud appears blurry because a time average of frames was used while illuminating the bubble cloud to get the image.
  • the cloud measures ⁇ 200um diameter at the narrowest point and ⁇ 300 urn along the longest axis.
  • the bubble cloud size increases and cavitation action becomes more aggressive as drive voltage and number of burst cycles increases. This can be used to control the speed and amount of tissue being ablated at a time.
  • results from subsequent measurements are shown in FIG. 13B.
  • the cavitation bubble cloud is shown as a white spot here due to averaging of multiple camera frames to produce an image.
  • This bubble cloud was generated in degassed, deionized water with a 6.8 MHz single-cycle, single-ended 173 V pulse at a 50 Hz repetition rate.
  • a 26 gauge needle, nominal diameter 464 ⁇ , can be seen in the image to provide scale for the bubble cloud which measures 124 ⁇ diameter with a length of 263 ⁇ .
  • FIG. 13B Both images demonstrate the ability to create a bubble cloud, however, FIG. 13B involved a longer time between pulses, and was only a single-cycle pulse. The smaller size of the bubble in FIG. 13B can be attributed to the lower number of cycles as this makes the overall volume which exceeds cavitation pressure smaller.
  • Imaging was performed using a Zeiss Discovery. V20 Stereo Microscope and a Zeiss Axiocam ERc 5s digital camera (Carl Zeiss Microscopy GmbH, Jena, Germany) where the bubble cloud was illuminated perpendicular to the direction of imaging, and multiple images were acquired and averaged to recreate the full bubble cloud shape.
  • This imaging is easy to do while cavitation occurs in water; however, in tissue, visual imaging would be impossible at-depth, so the tool presented here can instead be modified with the addition of a central hole, allowing an ultrasound probe through the center to image while ablation is being performed.
  • a co-registered imaging and ablation tool is described.
  • the focusing lens which was schematically illustrated in FIG. 1 , was modified to add a 4 mm x 4 mm hole through the center allowing any imaging tool to image the ablation area, as illustrated in FIG. 2.
  • the imaging device employed in the present example was developed in-house, and was fabricated as a 40 MHz, 64- element phased-array transducer packaged in a 2.5 x 3.1 mm endoscopic form factor. This endoscopic phased array was fully characterized by Bezanson et. al. in 2014 [A. Bezanson, R. Adamson, and J. A. Brown, "Fabrication and performance of a miniaturized 64- element high-frequency endoscopic phased array," IEEE Transactions on
  • FIG. 14 shows the example ultrasound endoscope with a machined lens positioned at the end as was used during co-registered ablation.
  • the device may be mounted at the end of a hand-held tool for fast user guidance so that ablation points can be targeted on-the-fly if desired.
  • the tight lens focus and small ablation spot-size, as well as the high-resolution endoscope imaging window renders the device as having potential for use in small animal studies where internal ablation, or highly targeted neural ablation with minimal tissue heating, may be desired.
  • the configuration shown in FIG. 14 provides one example orientation of the ablation lens relative to the endoscope.
  • the endoscope tip may be recessed from the lens curvature so that the endoscope does not occlude the ablation tool, while still having the ablation zone centered in the imaging window.
  • the lens-composite stack may be encased to ensure the composite remains air-backed. Preliminary testing of the co-registered device found that, with the current drive electronics and the missing lens area needed to accommodate the imaging probe, a higher voltage was needed to consistently cavitate.
  • FIG. 14 demonstrates this need for a higher drive voltage with a plot of pressure vs drive voltage for the transducer with a hole in the center.
  • FIGS. 15A and 15B show the peak-to-peak pressure, and the peak negative pressure, respectively, versus drive voltage for the transducer with a hole for co- registration for two different measurements. It is noted that the voltage needed to reach 10 MPa was ⁇ 2.5 times larger than for the non-co-registered transducer. This is related to the transducer hole reducing the ability of the transducer to focus.
  • a oneway response to a single cycle pulse at 6.8 MHz is shown in the top-left corner of FIG. 15A, where the response is good, however a tail in the response is seen.
  • FIGS. 16A and 16B the 2D pressure profile of the co-registered transducer is shown, with the -6 dB width shown in FIG. 16A, and the -3 dB width shown in FIG. 16B.
  • the radial beam width appears to be narrower compared to the transducer without a hole, however, there is more energy in the side lobes as seen by the higher pressures off center which is caused by the transducer hole reducing the ability of the lens to focus to a single point.
  • the pressure field shows visible side lobes at ⁇ 0.2 mm, 10 dB below the peak in the radial direction.
  • the transducer was driven with a 20 cycle pulse train at 6.8 MHz using a pulse-repetition frequency of 50 Hz.
  • pressure increased at a rate of 0.29 MPa / volt whereas for this co-registered imaging device the pressure increases at a rate of 0.1 MPa / volt.
  • FIGS. 17A and 17B histotripsy cavitation was performed in ex-vivo chinchilla cerebral tissue, the results of which are shown in FIGS. 17A and 17B.
  • the highly specular tissue is the cerebellum granular layer
  • the large dark regions are the molecular layer
  • thin dark tracts which are white matter.
  • the ability to identify these regions using ultrasound is important for targeting specific parts of the brain.
  • FIG. 17B histotripsy ablation is in progress, where the bright circular region between 7 mm and 8 mm depth is the bubble cloud in the process of ablating.
  • ablation is performed by driving the transducer with a 6.8 MHZ, 400 V, single-ended 10 cycle pulse train at a pulse repetition frequency (PRF) of 1000 Hz.
  • PRF pulse repetition frequency
  • Above the bubble cloud a channel can be seen from the cerebellum surface down to the ablation zone, where cavitation was initiated at the surface and then plunged into the tissue.
  • the bright streaks in the image are electronic noise from the histotripsy pulser and can be removed by synchronizing the histotripsy pulses to occur between image lines.
  • peak pressure should be less than three times the shock scattering threshold in the treated tissue.
  • the -3dB radial beam width is measured at 0.1 16 mm and the focal length, or -3dB axial beam width, is
  • a small, 10 mm aperture imaging and ablation device can create a histotripsy bubble cloud capable of ablating tissue with a focal zone and imaging capability allowing potentially sub-millimeter ablation accuracy.
  • the simplicity of the present example design should facilitate the creation of multiple tools without significant cost, while the demonstration of cavitation in water at a drive voltage of 173 V suggests driving the tool directly without matching circuitry or a transformer can keep the cost of the drive electronics low as well.
  • Side lobes on the non-co-registered device and on the co-registered device are close to -20 dB and -10 dB, respectively, suggesting that the probability of cavitation outside the focus is low.
  • a measured 59% one-way bandwidth allows the device to operate with a single-cycle or two-cycle pulse, maintaining a tight focus.
  • the present example device and the preceding example embodiments may be employed to provide an endoscopic imaging and ablation histotripsy tool.
  • Simulations were performed to investigate, for devices having an intermediate layer, the dependence of device performance on the acoustic impedance of the intermediate layer. Simulations were performed for devices having the same properties as in the previously described simulations, with an intermediate layer thickness of 20 microns, and with different values of the acoustic impedance of the intermediate layer. The simulations revealed that if the acoustic impedance of the intermediate layer is lower than that of both the piezoelectric layer and the acoustic lens, then a resonant increase in power is observed in the acoustic frequency spectrum.
  • the acoustic impedance of the intermediate layer is equal to the lowest of the acoustic impedances of the acoustic lens and the piezoelectric layer, then the resonance behavior of the acoustic power spectrum with resonant peaks having increased power output is not observed. If the acoustic impedance of the intermediate layer is between that of the piezoelectric layer and acoustic lens, then a resonant increase in power is also not observed. Furthermore, if the acoustic impedance of the intermediate layer is higher than that of both the piezoelectric layer and the acoustic lens, then a slight decrease in power is observed at higher frequencies.

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Abstract

L'invention concerne des systèmes et dispositifs permettant de générer des impulsions ultrasonores focalisées sur la base d'un ensemble transducteur comprenant une couche piézoélectrique couplée à une lentille acoustique. Dans certains modes de réalisation donnés à titre d'exemple, la couche piézoélectrique est un matériau piézoélectrique composite dont l'impédance acoustique est configurée pour correspondre à l'impédance acoustique de la lentille acoustique. La lentille acoustique peut être formée à partir d'aluminium ou d'un alliage de ce dernier, et peut présenter une surface distale à profil non sphérique afin de produire une région focale plus petite qu'une lentille sphérique équivalente. La lentille acoustique peut avoir une ouverture inférieure à l'unité. Dans certains modes de réalisation, la lentille acoustique est revêtue d'une couche d'adaptation d'impédance acoustique polymère qui est compatible avec un dépôt par dépôt chimique en phase vapeur, tel qu'un polymère à base de p-xylylène. Dans certains modes de réalisation, la lentille acoustique est formée à partir d'aluminium ou d'un alliage de ce dernier, et la couche d'adaptation d'impédance acoustique polymère est une couche de parylène.
EP18849828.1A 2017-09-01 2018-08-30 Ensemble transducteur pour générer des ultrasons focalisés Pending EP3675959A4 (fr)

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